U.S. patent number 10,667,750 [Application Number 16/060,423] was granted by the patent office on 2020-06-02 for method and system for sensing by modified nanostructure.
This patent grant is currently assigned to Ramot at Tel-Aviv University Ltd.. The grantee listed for this patent is Ramot at Tel-Aviv University Ltd.. Invention is credited to Vadim Krivitsky, Fernando Patolsky, Marina Zverzhinetsky.
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United States Patent |
10,667,750 |
Patolsky , et al. |
June 2, 2020 |
Method and system for sensing by modified nanostructure
Abstract
A method of detecting a presence and/or concentration of a
marker, e.g., a marker, in a liquid, e.g., a liquid, is disclosed.
The method comprises: contacting the liquid with a sensor having an
immobilized affinity moiety interacting with the marker and being
configured to generate a detectable signal responsively to the
interaction. The method further comprises washing the liquid off
the sensor, and detecting the presence and/or concentration of the
marker based on a detectable signal received from the sensor within
a time-window beginning a predetermined time period after a
beginning time of the washing.
Inventors: |
Patolsky; Fernando (Rehovot,
IL), Krivitsky; Vadim (Bney-Ayish, IL),
Zverzhinetsky; Marina (Rishon-LeZion, IL) |
Applicant: |
Name |
City |
State |
Country |
Type |
Ramot at Tel-Aviv University Ltd. |
Tel-Aviv |
N/A |
IL |
|
|
Assignee: |
Ramot at Tel-Aviv University
Ltd. (Tel-Aviv, IL)
|
Family
ID: |
59012779 |
Appl.
No.: |
16/060,423 |
Filed: |
December 8, 2016 |
PCT
Filed: |
December 08, 2016 |
PCT No.: |
PCT/IL2016/051319 |
371(c)(1),(2),(4) Date: |
June 08, 2018 |
PCT
Pub. No.: |
WO2017/098517 |
PCT
Pub. Date: |
June 15, 2017 |
Prior Publication Data
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Document
Identifier |
Publication Date |
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US 20180372678 A1 |
Dec 27, 2018 |
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Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
Issue Date |
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62264944 |
Dec 9, 2015 |
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62264913 |
Dec 9, 2015 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
G01N
33/5005 (20130101); H01L 23/532 (20130101); G01N
27/4146 (20130101); A61B 5/6847 (20130101); A61B
5/1473 (20130101); H01L 29/16 (20130101); A61B
5/4866 (20130101); H01L 29/06 (20130101); H01L
29/0673 (20130101); A61B 5/6833 (20130101); H01L
51/0545 (20130101); B82Y 15/00 (20130101); H01L
51/0049 (20130101); H01L 51/0558 (20130101) |
Current International
Class: |
A61B
5/00 (20060101); G01N 33/50 (20060101); G01N
27/414 (20060101); H01L 23/532 (20060101); H01L
29/06 (20060101); A61B 5/1473 (20060101); H01L
29/16 (20060101); B82Y 15/00 (20110101); H01L
51/05 (20060101); H01L 51/00 (20060101) |
References Cited
[Referenced By]
U.S. Patent Documents
Foreign Patent Documents
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1669748 |
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Jun 2006 |
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EP |
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1806414 |
|
Jul 2007 |
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EP |
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2009-540798 |
|
Nov 2009 |
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JP |
|
2010-515887 |
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May 2010 |
|
JP |
|
2012-511156 |
|
May 2012 |
|
JP |
|
WO 2004/034025 |
|
Apr 2004 |
|
WO |
|
WO 2005/004204 |
|
Jan 2005 |
|
WO |
|
WO 2008/027078 |
|
Mar 2008 |
|
WO |
|
WO 2008/030395 |
|
Mar 2008 |
|
WO |
|
WO 2008/083446 |
|
Jul 2008 |
|
WO |
|
WO 2009/104180 |
|
Aug 2009 |
|
WO |
|
WO 2010/115143 |
|
Oct 2010 |
|
WO |
|
WO 2011/000443 |
|
Jan 2011 |
|
WO |
|
WO 2012/082494 |
|
Jun 2012 |
|
WO |
|
WO 2012/137207 |
|
Oct 2012 |
|
WO |
|
WO 2015/059704 |
|
Apr 2015 |
|
WO |
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WO 2017/098517 |
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Jun 2017 |
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WO |
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WO 2017/098518 |
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Jun 2017 |
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WO |
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Other References
Notice of Reasons for Rejection dated Oct. 9, 2018 From the Japan
Patent Office Re. Application No. 2016-525073 and Its Translation
Into English. (10 Pages). cited by applicant .
Communication Pursuant to Article 94(3) EPC dated May 9, 2017 From
the European Patent Office Re. Application No. 14796555.2. (6
Pages). cited by applicant .
Communication Pursuant to Article 94(3) EPC dated Feb. 20, 2018
From the European Patent Office Re. Application No. 14796555.2. (6
Pages). cited by applicant .
International Preliminary Report on Patentability dated May 6, 2016
From the International Bureau of WIPO Re. Application No.
PCT/IL2014/050921. cited by applicant .
International Search Report and the Written Opinion dated Feb. 23,
2015 From the International Searching Authority Re. Application No.
PCT/IL2014/050921. cited by applicant .
International Search Report and the Written Opinion dated Mar. 26,
2017 From the International Searching Authority Re. Application No.
PCT/IL2016/051319. (12 Pages). cited by applicant .
International Search Report and the Written Opinion dated Mar. 29,
2017 From the International Searching Authority Re. Application No.
PCT/IL2016/051320. (12 Pages). cited by applicant .
Notice of Eligibility for Grant and Examination Report dated Jun.
11, 2018 From the Intellectual Property Office of Sinagpore, IPOS
Re. Application No. 11201602976X. (7 Pages). cited by applicant
.
Official Action dated Jul. 5, 2018 From the US Patent and Trademark
Office Re. U.S. Appl. No. 15/030,886. (54 pages). cited by
applicant .
Search Report and Written Opinion dated Sep. 21, 2016 From the
Intellectual Property Office of Singapore Re. Application No.
11201602976X. cited by applicant .
Written Opinion dated Sep. 6, 2017 From the Intellectual Property
Office of Singapore, IPOS Re. Application No. 11201602976X. (8
Pages). cited by applicant .
Chen et al. "Label-Free Cytokine Micro- and Nano-Biosensing Towards
Personalized Medicine of Systemic Inflammatory Disorders", Advanced
Drug Delivery Reviews, 95: 90-103, Available Online Sep. 15, 2015.
p .4, r-h Col., 1st Para, Fig.3. cited by applicant .
Chen et al. "Silicon Nanowire Filed-Effect Transistor-Based
Biosensors for Biomedical Diagnosis and Cellular Recording
Investigation", Nano Today, 6(2): 131-154, Available Online Mar. 8,
2011. cited by applicant .
Clavaguera et al. "Sup-PPM Detection of Nerve Agents Using
Chemically Functionalized Silicon Nanoribbon Field-Effect
Transistors", Angewandte Chemie International Edition, 49(24):
4063-4066, Jun. 2010. cited by applicant .
Cui et al. "Nanowire Nanosensors for Highly Sensitive and Selective
Detection of Biological and Chemical Species", Science, 293(5533):
1289-1292, Aug. 17, 2001. cited by applicant .
De et al. "Integrated Label-Free Silicon Naonowire Sensor Arrays
for (Bio)Chemical Analysis", The Analyst, XP055168035, 138(11):
3221-3229, Jan. 2013. Fig.2, figs.1-3, Abstract, p. 4, 2nd Col.
cited by applicant .
Duan et al. "Intracellular Recordings of Action Potentials by an
Extracellular Nanoscale Field-Effect Transistor", Nature
Nanotechnology, 7(3): 174-179, Published Online Dec. 18, 2011.
cited by applicant .
Garcia et al. "Enhanced Determination of Glucose by Microchip
Electrophoresis With Pulsed Amperometric Detection", Analytica
Chimica Acta, 508(1): 1-9, Apr. 15, 2004. cited by applicant .
Garcia-Perez et al. "Metabolic Fingerprinting With Capillary
Electrophoresis", Journal of Chromatography A, 1204(2): 130-139,
Available Online Jul. 12, 2008. cited by applicant .
Griffin et al. "Metabolic Profiles of Cancer Cells", Nature Reviews
Cancer, 4(7): 551-561, Jul. 2004. cited by applicant .
Holcomb et al. "Electrode Array Detector for Microchip Capillary
Electrophoresis", The Analyst, 134(3): 486-492, Published Online
Dec. 3, 2008. cited by applicant .
Hsiung et al. "Multiplex Reatl-Time Monitoring of Cellular
Metabolic Activity Using a Redox-Reactive Nanowire Biosensor", 17th
International Conference on Miniaturized Systems for Chemistry and
Life Sciences, Freiburg, Germany, Oct. 27-31, 2013, XP055167304, p.
1959-1961, Oct. 2013. cited by applicant .
Huang et al. "Real-Time and Label-Free Detection of the
Prostate-Specific Antigen in Human Serum by A Polycrystalline
Silicon Nanowire Field-Effect Transistor Biosensor", Analytical
Chemistry, 85(16): 7912-7918, Jul. 11, 2013. p. 7914, r-h Col., 1st
Para, Figs.3b, 4b, 6. cited by applicant .
Jeykumari et al. "Covalent Modification of Multiwalled Carbon
Nanotubes with Neutral Red for the Fabrication of an Amperometric
Hydrogen Peroxide Sensor", Nanotechnology 18(125501): 1-10, 2007.
cited by applicant .
Kleps et al. "Investigation of Silver-, Meso- and Nanoporous
Silicon Composite Layers for Biomedical Applications", Romanian
Journal of Information Science and Technology, 10(1): 97-111, 2007.
cited by applicant .
Kosaka et al. "Detection of Cancer Biomarkers in Serum Using a
Hybrid Mechanical and Optoplasmonic Nanosensor", Nature
Nanotechnology, 9(12): 1047-1053, Published Online Nov. 2, 2014.
cited by applicant .
Kraly et al. "Review: Microfluidic Applications in Metabolomics and
Metabolic Profiling", Analytica Chimica Acta, 653(1): 23-35,
Available Online Sep. 1, 2009. cited by applicant .
Krivitsky et al. "Antigen-Dissociation From Antibody-Modified
Nanotransistor Sensor Arrays as a Direct Biomarker Detection Method
in Unprocessed Biosamples", Nano Letters, 16(10): 6272-6281, Aug.
31, 2016. cited by applicant .
Krivitsky et al. "Si Nanowires Forest-Based On-Chip Biomolecular
Filtering, Separation and Preconcentration Devices: Nanowires Do It
All", Nano Letters, 12(9): 4748-4756, Aug. 2, 2012. cited by
applicant .
Li et al. "Sequence-Specific Label-Free DNA Sensors Based on
Silicon Nanowires", Nano Letters, 4(2): 245-247, Published on Web
Jan. 8, 2004. cited by applicant .
Lin et al. "Microscale LC-MS-NMR Platform Applied to the
Identification of Active Cyanobacterial Metabolites", Analytical
Chemistry, 80(21): 8045-8054, Nov. 1, 2008. cited by applicant
.
Lu et al. "A Nano-Ni Based Ultrasensitive Nonenzymatic
Electrochemical Sensor for Glucose: Enhancing Sensitivity Through a
Nanowire Array Strategy", Biosensors and Bioelectronics, 25(1):
218-223, Published Online Jul. 7, 2009. cited by applicant .
Lu et al. "Enzyme-Functionalized Gold Nanowires for the Fabrication
of Biosensors", Bioelectrochemistry, 71(2): 211-216, Published
Online Jun. 14, 2007. cited by applicant .
Lu et al. "Label-Free and Rapid Electrical Detection of hTSH With
CMOS-Compatible Silicon Nanowire Transistor Arrays", ACS Applied
Materials & Interfaces, 6(22): 20378-20384, Oct. 22, 2014.
Figs.4a, 4b, Table 1, p. 20381, 1-h Col., Last Para. cited by
applicant .
Marx "Tracking Metastasis and Tricking Cancer", Nature, 494(7435):
131-136, Feb. 7, 2013. cited by applicant .
McAlpine et al. "Highly Ordered Nanowire Arrays on Plastic
Substrates for Ultrasensitive Flexible Chemical Sensors", Nature
Materials, 6(5): 379-384, May 2007. cited by applicant .
Mohanty et al. "Field Effect Transistor Nanosensor for Breast
Cancer Diagnostics", ArXiv Preprint ArXiv, 1401.1168: 1-25, Jan. 6,
2014. p. 5, Section B, p. 10, 2nd Para, p. 14, 3rd Para. cited by
applicant .
Mu et al. "Silicon Nanowire Field-Effect Transistors--A Versatile
Class of Potentiometric Nanobiosensors", IEEE Access, 3: 287-302,
Apr. 22, 2015. p. 293, 1-h Col., 4th Para, p. 290, r-h Col., 1st
Para. cited by applicant .
Munoz-Pinedo et al. "Cancer Metabolism: Current Perspectives and
Future Directions", Cell Death and Disease, 3(1): e248-1-e248-10,
Published Online Jan. 12, 2012. cited by applicant .
Noor et al. "Silicon Nanowires as Field-Effect Transducers for
Biosensor Development: A Review", Analytica Chimica Acta, 825:
1-25, Available Online May 15, 2014. cited by applicant .
Northen et al. "Clathrate Nanostructures for Mass Spectrometry",
Nature, 449(7165): 1033-1037, Oct. 25, 2007. cited by applicant
.
Patolsky "Nanotechnology Tools in Biology and Medicine
Applications", YouTube [Online], Presentation, Summer School on
Nanomedicine and Innovation, The Marian Gertner Institute for
Medical Nanosystems, Raymond and Beverly Sackler School of
Chemistry, Tel Aviv University, Israel, Jun. 19, 2014. Video:
45:13-49:10 (mm:ss). cited by applicant .
Patolsky et al. "Electrical Detection of Single Viruses", Proc.
Natl. Acad. Sci. USA, PNAS, 101(39): 14017-14022, Sep. 28, 2004.
cited by applicant .
Patolsky et al. "Fabrication of Silicon Nanowire Devices for
Ultrasensitive, Label- Free, Real-Time Detection of Biological and
Chemical Species", Nature Protocols, 1(4): 1711-1724, Published
Online Nov. 16, 2006. cited by applicant .
Patolsky et al. "Nanowire-Based Biosensors", Analytical Chemistry,
78(13): 4260-4269, Jul. 1, 2006. cited by applicant .
Peretz-Soroka et al. "Optically-Gated Self-Calibrating Nanosensors:
Monitoring pH and Metabolic Activity of Living Cells", Nano
Letters, 13(7): 3157-3168, Jun. 17, 2013. cited by applicant .
Ramgir et al. "Nanowire-Based Sensors", Small, 6(16): 1705-1722,
Aug. 17, 2010. cited by applicant .
Shaijumon et al. "Catalytic Growth of Carbon Nanotubes Over Ni/Cr
Hydrotalcite-Type Anionic Clay and Their Hydrogen Storage
Properties", Applied Surface Science, 242: 192-198, 2005. cited by
applicant .
Shao et al. "Silicon Nanowire Sensors for Bioanalytical
Applications: Glucose and Hydrogen Peroxide Detection", Advanced
Functional Materials, 15(9): 1478-1482, Sep. 2005. cited by
applicant .
Shulaev "Metabolomics Technology and Bioinformatics", Briefings in
Bioinformatics, 7(2): 128-139, May 18, 2006. cited by applicant
.
Stern et al. "Label-Free Biomarker Detection From Whole Blood",
Nature Nanotechnology, 5(2): 138-142, Published Online Dec. 13,
2009. cited by applicant .
Stern et al. "Semiconducting Nanowire Field-Effect Transistor
Biomolecular Sensors", IEEE Transactions on Electron Devices,
55(11): 3119-3130, Nov. 2008. cited by applicant .
Su et al. "A Silicon Nanowire-Based Electrochemical Sensor With
High Sensitivity and Electrocatalytic Activity", Particle Particle
Systems Characterization, 30(4): 326-331, Apr. 2013. cited by
applicant .
Telg et al. "G- and G+ in the Raman Spectrum of Isolated Nanotube:
A Study on Resonance Conditions and Lineshape", Physica Status
Solidi (b), 245(10): 2189-2192, 2008. cited by applicant .
Timko et al. "Electrical Recording From Hearts With Flexible
Nanowire Device Arrays", Nano Letters, 9(2): 914-918, Published on
Web Jan. 26, 2009. cited by applicant .
Tyagi et al. "Patternable Nanowire Sensors for Electrochemical
Recording of Dopamine", Analytical Chemistry, 81(24): 9979-9984,
Dec. 15, 2009. cited by applicant .
Vlckova et al. "Determination of Cationic Neutrotransmitters and
Metabolites in Brain Homogenates by Microchip Electrophoresis and
Carbon Nanotube-Modified Amperometry", Journal of Chromatography A,
1142(2): 214-221, 2007. cited by applicant .
Wanekeya et al. "Nanowire-Based Electrochemical Biosensors",
Electroanalysis, XP055167317, 18(6): 533-550, Mar. 1, 2006. cited
by applicant .
Wang et al. "A NEMS Thermal Biosensor for Metabolic Monitoring
Applications", Journal of Microelectromechanical Systems, 17(2):
318-327, Apr. 2008. cited by applicant .
Wang et al. "Simultaneous Microchip Enzymatic Measurements of Blood
Lactate and Glucose", Analytica Chimica Acta, 585(1): 11-16,
Published Online Dec. 9, 2006. cited by applicant .
Yang et al. "Gold Nanoparticle Modified Silicon Nanowires as
Biosensors", Nanotechnology, 17(11): S276-S279, May 19, 2006. cited
by applicant .
Yin et al. "A Hydrogen Peroxide Electrocheinical Sensor Based on
Silver Nanopartides Decorated Silicon Nanowire Arrays"
Electrochimica Acta, 56: 3884-3889, 2011. cited by applicant .
Yun et al. "On-Line Carbon Nanotube-Based Biosensors in
Microfluidic Channels", Nanosensors, Microsensors, and Biosensors
and Systems, Proceedings of the SPIE, XP055167836, 6528:
65280-1-65280-10, Apr. 4, 2007. Abstract, Figs.3-5. cited by
applicant .
Zayats et al. "An Integrated NAD+-Dependent Enzyme-Functionalized
Field-Effect Transistor (ENFET) System: Development of a Lactate
Biosensor", Biosensor & Bioelectronics, XP055450950, 15(11-12):
671-680, Dec. 1, 2000. cited by applicant .
Zheng et al. "Multiplexed Electrical Detection of Cancer Markers
With Nanowire Sensor Arrays", Nature Biotechnology, 23(10):
1294-1301, Oct. 2005. cited by applicant .
Supplementary European Search Report and the Euorpan Search Opinion
dated Jul. 11, 2019 From the European Patent Office Re. Application
No. 16872559.6. (10 Pages). cited by applicant .
Supplementary European Search Report and the European Search
Opinion dated Jul. 10, 2019 From the European Patent Office Re.
Application No. 16872558.8. (12 Pages). cited by applicant .
Elnathan et al. "Biorecognition Layer Engineering: Overcoming
Screening Limitations of Nanowire-Based FET Devices", Nano Letters,
XP055366864, 12(10): 5245-5254, Published Online Sep. 10, 2012.
cited by applicant .
Gao et al. "General Strategy for Biodetection in High Ionic
Strength Solutions Using Transistor-Based Nanoelectronic Sensors",
Nano Letters, XP055317106, 15(3): 2143-2148, Published Online Feb.
9, 2015. cited by applicant .
Hwang et al. "Biodegradable Elastomers and Silicon
Nanomembranes/Nanoribbons for Stretchable, Transient Electronics,
and Biosensors", Nano Letters, XP055601726, 15(5): 2801-2808,
Published Online Feb. 23, 2015. cited by applicant .
Kim et al. "Direct Label-Free Electrical Immunodetection in Human
Serum Using a Flow-Through-Apparatus Approach With Integrated
Field-Effect Transistors", Biosensors and Bioelectronics,
XP029490340, 25(7): 1767-1773, Available Online Dec. 29, 2009.
cited by applicant .
Rajan et al. "Performance Limitations for Nanowire/Nanoribbon
Biosensors", Wiley Interdisciplinary Reviews: Nanomedicine and
Nanobiotechnology, XP055593106, 5(6): 629-645, Published Online
Jul. 29, 2013. cited by applicant .
Ramachandran et al. "A Rapid, Multiplexed, High-Throughput
Flow-Through Membrane Immunoassay: A Convenient Alternative to
ELISA", Diagnostics, XP055549823, 3(2): 244-260, Published Online
Apr. 2, 2013. cited by applicant .
Official Action dated Mar. 27, 2020 from the US Patent and
Trademark Office Re. U.S. Appl. No. 16/297,665. (41 pages). cited
by applicant.
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Primary Examiner: Reames; Matthew L
Parent Case Text
RELATED APPLICATIONS
This application is a National Phase of PCT Patent Application No.
PCT/IL2016/051319 having International filing date of Dec. 8, 2016,
which claims the benefit of priority under 35 USC .sctn. 119(e) of
U.S. Provisional Patent Applications Nos. 62/264,913 and
62/264,944, both filed on Dec. 9, 2015. The contents of the above
applications are all incorporated by reference as if fully set
forth herein in their entirety.
Claims
What is claimed is:
1. A method of detecting a presence and/or concentration of a
marker in a liquid, the method comprising: contacting the liquid
with a sensor having an immobilized affinity moiety interacting
with the marker and being configured to generate a detectable
signal responsively to said interaction, said interaction being
characterized by a K.sub.D which is equal or less than 10.sup.-5 M;
washing said liquid off said sensor; and detecting the presence
and/or concentration of the marker based on a detectable signal
received from said sensor within a time-window beginning after a
beginning time of said washing, said detectable signal being
indicative of desorption kinetic of the marker, wherein said
detecting is based on said desorption kinetic.
2. The method of claim 1, wherein said marker is a biomarker.
3. The method according to claim 1, wherein said liquid is a
biological liquid.
4. The method according to claim 1, comprising monitoring said
detectable signal also before said beginning time of the
time-window, but said detecting the presence and/or concentration
of the marker is not based on any signal received from the sensor
before said beginning time of the time-window.
5. The method according to claim 1, wherein said time-window begins
at least 30 seconds after said beginning time of said washing.
6. The method according to claim 1, further comprising monitoring
said detectable signal from said beginning time of said washing,
and identifying said beginning of said time-window based on a
change in a time-dependence of said signal.
7. The method according to claim 6, wherein said beginning of said
time-window is defined at a time point, which is after said
beginning time of said washing, and at which a rate of change of
said signal, in absolute value, is below a predetermined
threshold.
8. The method according to claim 1, wherein said affinity moiety
comprises an immunogenic moiety.
9. The method according to claim 1, wherein said affinity moiety
comprises an immunogenic moiety, which comprises an antibody or a
fragment thereof.
10. The method according to claim 1, wherein said affinity moiety
comprises an immunogenic moiety, which comprises an antigen, and
wherein said marker is a biomarker which comprises an antibody to
said antigen.
11. The method according to claim 1, wherein said affinity moiety
comprises a ligand and said marker is a biomarker which comprises a
receptor.
12. The method according to claim 1, wherein said sensor is a
nanostructure and the affinity moiety is immobilized on a surface
of said nanostructure, wherein said contacting the liquid comprises
introducing the liquid to a sensing chamber containing therein said
sensor and an additional sensor which is also a nanostructure but
is devoid of the affinity moiety, wherein said washing comprises
washing said sensing chamber, wherein the method comprises
comparing a rate of returning to a baseline of said detectable
signal with a rate of returning to a baseline of a background
signal received from said additional sensor, and wherein said
detection of the presence and/or concentration of the marker is
based on said comparison.
13. The method according to claim 1, wherein said sensor is a
nanostructure and the affinity moiety is immobilized on a surface
of said nanostructure.
14. The method according to claim 1, wherein said sensor is a
transistor.
15. The method according to claim 1, wherein said sensor is a
transistor, having a nanostructure as a channel and wherein the
affinity moiety is immobilized on a surface of said
nanostructure.
16. The method according to claim 14, wherein said transistor is a
field-effect transistor.
Description
FIELD AND BACKGROUND OF THE INVENTION
The present invention, in some embodiments thereof, relates to
sensing and, more particularly, but not exclusively, to a methods
and system for detecting a marker, such as, but not limited to, a
biomarker, in a liquid, such as, but not limited to, biological
liquid.
The development of efficient bio-molecular separation and
purification techniques is of high importance in modern genomics,
proteomics, and bio-sensing areas, primarily due to the fact that
most bio-samples are mixtures of high diversity and complexity.
Most of the currently-practiced techniques lack the capability to
rapidly and selectively separate and concentrate specific target
proteins from a complex bio-sample, and are difficult to integrate
with lab-on-a-chip sensing devices.
Semiconducting nanowires are known to be extremely sensitive to
chemical species adsorbed on their surfaces. For a nanowire device,
the binding of a charged analyte the surface of the nanowire leads
to a conductance change, or a change in current flowing through the
wires. The 1D (one dimensional) nanoscale morphology and the
extremely high surface-to-volume ratio make this conductance change
to be much greater for nanowire-based sensors versus planar FETs
(field-effect transistors), increasing the sensitivity to a point
that single molecule detection is possible.
Nanowire-based field-effect transistors (NW-FETs) have therefore
been recognized in the past decade as powerful potential new
sensors for the detection of chemical and biological species. See,
for example, Patolsky et al., Analytical Chemistry 78, 4260-4269
(2006); Stern et al., IEEE Transactions on Electron Devices 55,
3119-3130 (2008); Cui et al., Science 293, 1289-1292 (2001);
Patolsky et al. Proceedings of the National Academy of Sciences of
the United States of America 101, 14017-14022 (2004), all being
incorporated by reference as if fully set forth herein.
Studies have also been conducted with nanowire electrical devices
for the simultaneous multiplexed detection of multiple biomolecular
species of medical diagnostic relevance, such as DNA and proteins
[Zheng et al., Nature Biotechnology 23, 1294-1301 (2005); Timko et
al., Nano Lett. 9, 914-918 (2009); Li et al., Nano Lett. 4, 245-247
(2004)].
Generally, in a NW-FET configuration, the gate potential controls
the channel conductance for a given source drain voltage (VSD), and
modulation of the gate voltage (VGD) changes the measured
source-drain current (ISD). For NW sensors operated as FETs, the
sensing mechanism is the field-gating effect of charged molecules
on the carrier conduction inside the NW. Compared to devices made
of micro-sized materials or bulk materials, the enhanced
sensitivity of nanodevices is closely related to the reduced
dimensions and larger surface/volume ratio. Since most of the
biological analyte molecules have intrinsic charges, binding on the
nanowire surface can serve as a molecular gate on the
semiconducting SiNW [Cui et al., 2001, supra].
Antibody/enzyme nanowire FET devices which target metabolites via
binding affinity have been disclosed in, for example, Lu et al.
Bioelectrochemistry 2007, 71(2): 211-216; Patolsky et al.
Nanowire-based biosensors. Anal Chem 2006, 78(13): 4260-4269; and
Yang et al. Nanotechnology 2006, 17(11): S276-S279.
Electrochemically-sensitive nanowire sensors for detecting
metabolites by oxidative reactions have been disclosed in, for
example, Lu et al. Biosens Bioelectron 2009, 25(1): 218-223; Shao
et al. Adv Funct Mater 2005, 15(9): 1478-1482; Su et al. Part Part
Syst Char 2013, 30(4): 326-331; and Tyagi et al. Anal Chem 2009,
81(24): 9979-9984.
U.S. Pat. No. 7,619,290, U.S. Patent Application having publication
No. 2010/0022012, and corresponding applications, teach nanoscale
devices composed of, inter alia, functionalized nanowires, which
can be used as sensors.
Clavaguera et al. disclosed a method for sub-ppm detection of nerve
agents using chemically functionalized silicon nanoribbon
field-effect transistors [Clavaguera et al., Angew. Chem. Int. Ed.
2010, 49, 1-5].
SiO.sub.2 surface chemistries were used to construct a
`nano-electronic nose` library, which can distinguish acetone and
hexane vapors via distributed responses [Nature Materials Vol. 6,
2007, pp. 379-384].
U.S. Patent Application having Publication No. 2010/0325073
discloses nanodevices designed for absorbing gaseous NO. WO
2011/000443 describes nanodevices which utilize functionalized
nanowires for detecting nitro-containing compounds.
Duan et al. [Nature Nanotechnology, Vol. 7, 2012, pp. 174-179]
describes a silicon nanowire FET detector and an electrically
insulating SiO2 nanotube that connects the FET to the intracellular
fluid (the cytosol). When there is a change in transmembrane
potential Vm, the varying potential of the cytosol inside the
nanotube gives rise to a change in the conductance G of the
FET.
Kosaka et al. [Nature Nanotechnology, Vol. 9, 2014, pp. 1047-1053]
discloses detection of cancer biomarkers in serum using
surface-anchored antibody.
Krivitsky et al. [Nano letters 2012, 12(9): 4748-4756] describe an
on-chip all-SiNW filtering, selective separation, desalting, and
preconcentration platform for the direct analysis of whole blood
and other complex biosamples. The separation of required protein
analytes from raw bio-samples is first performed using a
antibody-modified roughness-controlled SiNWs forest of ultralarge
binding surface area, followed by the release of target proteins in
a controlled liquid media, and their subsequent detection by
SiNW-based FETs arrays fabricated on the same chip platform.
WO 2015/059704 discloses an integrated microfluidic nanostructure
sensing system, comprised of one or more sensing compartments
featuring a redox-reactive nanostructure FET array which is in
fluid communication with one or more sample chambers. This system
has been shown to perform multiplex real-time monitoring of
cellular metabolic activity in physiological solutions, and was
demonstrated as an efficient tool in promoting the understanding of
metabolic networks and requirements of cancers for personalized
medicine.
Additional background art includes, for example, Chen et al., Nano
Today (2011) 6, 131-54, and references cited therein; and Stern et
al., Nature Nanotechnology, 2009.
SUMMARY OF THE INVENTION
According to an aspect of some embodiments of the present invention
there is provided a method of detecting a presence and/or
concentration of a marker in a liquid. The method comprises:
contacting the liquid with a sensor having an immobilized affinity
moiety interacting with the marker and being configured to generate
a detectable signal responsively to the interaction. The method
also comprises washing the liquid off the sensor; and detecting the
presence and/or concentration of the marker based on a detectable
signal received from the sensor within a time-window beginning a
predetermined time period after a beginning time of the
washing.
According to some embodiments of the invention the marker is a
biomarker.
According to some embodiments of the invention the liquid is a
biological liquid.
According to some embodiments of the invention the detection is not
based on signal received from the sensor before the beginning time
of the time-window.
According to some embodiments of the invention the predetermined
time period is at least 30 seconds.
According to some embodiments of the invention the method comprises
monitoring the detectable signal from the beginning of the washing,
and identifying the beginning of the time-window based on a change
in a time-dependence of the signal.
According to some embodiments of the invention the affinity moiety
comprises an immunogenic moiety.
According to some embodiments of the invention the immunogenic
moiety comprises an antibody or a fragment thereof.
According to some embodiments of the invention the immunogenic
moiety comprises an antigen and wherein the marker is a biomarker
which comprises an antibody to the antigen.
According to some embodiments of the invention the affinity moiety
comprises a ligand and the marker is a biomarker which comprises a
receptor.
According to some embodiments of the invention the method comprises
contacting the liquid with a sensor which is non-specific to the
marker and washing the liquid also off the non-specific sensor,
wherein the detection of the presence and/or concentration of the
marker is based on a comparison between the detectable signal and a
background signal received from the non-specific sensor.
According to an aspect of some embodiments of the present invention
there is provided a system for detecting a presence and/or
concentration of a marker in a liquid. The system comprises a
fluidic device having a sensing chamber and a sensor in the sensing
chamber, the sensor having an immobilized affinity moiety
interacting with the marker and being configured to generate a
detectable signal responsively to the interaction; a flow control
system for introducing a washing buffer to the chamber to wash the
liquid off the sensor; and a signal analyzer for analyzing
detectable signal received from the sensor over a time-window
beginning a predetermined time period after a beginning time of the
washing, to detect the presence and/or concentration of the
marker.
In any of the embodiments, the interaction is optionally and
preferably characterized by a K.sub.D which is equal or less than
10.sup.-5 M or a K.sub.D which is equal or less than 10.sup.-6 M or
a K.sub.D which is equal or less than 10.sup.-7 M or a K.sub.D
which is equal or less than 10.sup.-8 M or a K.sub.D which is equal
or less than 10.sup.-9 M or a K.sub.D which is equal or less than
10.sup.-10 M.
According to some embodiments of the invention the signal analyzer
is configured to discard from the analysis signal received from the
sensor before the beginning time of the time-window.
According to some embodiments of the invention the signal analyzer
is configured for monitoring the detectable signal from the
beginning of the washing, and identifying the beginning of the
time-window based on a change in a time-dependence of the
signal.
According to some embodiments of the invention the beginning of the
time-window is defined at a time point at which a rate of change of
the signal, in absolute value, is below a predetermined
threshold.
According to some embodiments of the invention the fluidic device
further comprises a sensor that is non-specific the marker in the
sensing chamber, signal analyzer is configured for comparing
between the detectable signal and a background signal received from
the non-specific sensor, to detect the presence and/or
concentration of the marker based on the comparison.
According to some embodiments of the invention the sensor is a
nanostructure and the affinity moiety is immobilized on a surface
of the nanostructure.
According to some embodiments of the invention the sensor is a
transistor.
According to some embodiments of the invention the sensor is a
transistor, having a nanostructure as a channel and wherein the
affinity moiety is immobilized on a surface of the
nanostructure.
According to some embodiments of the invention the transistor is a
field-effect transistor.
BRIEF DESCRIPTION OF SEVERAL VIEWS OF THE DRAWINGS
The patent or application file contains at least one drawing
executed in color. Copies of this patent or patent application
publication with color drawing(s) will be provided by the Office
upon request and payment of the necessary fee.
Some embodiments of the invention are herein described, by way of
example only, with reference to the accompanying drawings. With
specific reference now to the drawings in detail, it is stressed
that the particulars shown are by way of example and for purposes
of illustrative discussion of embodiments of the invention. In this
regard, the description taken with the drawings makes apparent to
those skilled in the art how embodiments of the invention may be
practiced.
In the drawings:
FIGS. 1A-G are images (FIGS. 1A and 1B) and schematic illustrations
of a fabrication process (FIGS. 1C-G) of a 20 nm diameter P-type
silicon nanowires (SiNW) FET device, according to some embodiments
of the present invention;
FIG. 2 is a schematic illustration of a process for the
modification of SiNW FET device with antibodies, according to some
embodiments of the present invention;
FIG. 3 shows an exemplary response calibration curve of an
antibody-modified SiNW FET, upon introducing of an
antigen-containing buffer, following by introducing a washing
buffer into a fluidic system containing the a plurality of SiNW
FET, as obtained in experiments performed according to some
embodiments of the present invention;
FIG. 4 shows a calibrated response of an antibody-modified SiNW
FET, in a bio-sample, as obtained in experiments performed
according to some embodiments of the present invention;
FIG. 5 shows a response calibration curve in the presence of the
bio-sample of FIG. 4, but after initiation of a washing operation,
according to some embodiments of the present invention;
FIG. 6 shows two response calibration curves obtainable in the
presence of a bio-sample, after the initiation of a washing
operation, as obtained in experiments performed according to some
embodiments of the present invention;
FIG. 7 is a flowchart diagram schematically illustrating of a
method suitable for detecting a presence and/or concentration of a
biomarker in a biological liquid, according to some embodiments of
the present invention;
FIGS. 8A and 8B are schematic illustrations of sensors suitable for
use according to some embodiments of the present invention;
FIG. 9 is a schematic illustration of a system for detecting a
presence and/or concentration of a marker, e.g., biomarker in a
liquid e.g., biological liquid, according to some embodiments of
the present invention;
FIG. 10 is a schematic illustration describing an operation
principle of a detection of an antigen from antibody-modified SiNW
FET sensing devices, according to some embodiments of the present
invention;
FIG. 11 shows normalized association response of a representative
anti-CA 15-3 immobilized SiNW FET sensing device against various
concentrations of the target antigen CA 15-3 in unprocessed serum
samples, as obtained during experiments performed according to some
embodiments of the present invention;
FIG. 12 shows normalized electrical response of a representative
anti-troponin antibody-modified SiNW FET sensing device to the
association and dissociation of its specific antigen troponin T
under low-ionic strength conditions, as obtained during experiments
performed according to some embodiments of the present
invention;
FIG. 13 shows anti-troponin antibody dissociation curve in serum,
together with a fit, under slow-flow conditions, as obtained during
experiments performed according to some embodiments of the present
invention;
FIGS. 14A and 14B show results of a comparison between dissociation
kinetics of troponin T antigen-containing serum sample and a
control troponin T-free serum sample, as obtained during
experiments performed according to some embodiments of the present
invention;
FIG. 15 shows the serum dissociation curve, together with a fit,
under slow-flow conditions, as obtained during experiments
performed according to some embodiments of the present
invention;
FIG. 16 shows anti-CA 15-3 antibody dissociation curve in serum,
together with a fit, under slow-flow conditions, as obtained during
experiments performed according to some embodiments of the present
invention;
FIGS. 17A-D show concentration-dependent sensing of CA 15-3 antigen
in unprocessed serum samples, at a flow rate of 1 chamber-volume
exchange per minute, as obtained during experiments performed
according to some embodiments of the present invention;
FIG. 18 shows regeneration curve of a CA 15-3 antigen from its
antibody-modified nanowire device, as obtained during experiments
performed according to some embodiments of the present
invention;
FIGS. 19A and 19B demonstrate multiplex single-chip differential
detection of the CA 15-3 antigen using specific and nonspecific
chemically modified SiNWs FET devices, at a flow rate of 330
chamber-volumes per minute, as obtained during experiments
performed according to some embodiments of the present
invention;
FIG. 20 shows antigen-free serum dissociation curve under fast-flow
conditions, together with a fit, as obtained during experiments
performed according to some embodiments of the present
invention;
FIG. 21 shows antigen-free untreated blood dissociation curve under
slow-flow conditions, together with a fit, as obtained during
experiments performed according to some embodiments of the present
invention;
FIGS. 22A and 22B show electrical characterization of a p-type SiNW
FET nanodevices under water-gate configuration, as obtained during
experiments performed according to some embodiments of the present
invention; and
FIGS. 23A and 23B are Scanning Electron Microscope (SEM) images of
the SiNWs obtained during experiments performed according to some
embodiments of the present invention.
DESCRIPTION OF SPECIFIC EMBODIMENTS OF THE INVENTION
The present invention, in some embodiments thereof, relates to
sensing and, more particularly, but not exclusively, to a methods
and system for detecting a marker, such as, but not limited to, a
biomarker, in a liquid, such as, but not limited to, biological
liquid. Before explaining at least one embodiment of the invention
in detail, it is to be understood that the invention is not
necessarily limited in its application to the details of
construction and the arrangement of the components and/or methods
set forth in the following description and/or illustrated in the
drawings and/or the Examples. The invention is capable of other
embodiments or of being practiced or carried out in various
ways.
The present inventors have designed a sensing system and method,
which is usable for sensing and optionally and preferably
monitoring the presence, and more preferably amount, of a marker in
a liquid. The sensing system and method can be used for multiplex
real-time monitoring of many types of markers in many types of
liquids.
The sensing system of the present embodiments can be used in many
applications, including without limitation, chemical applications,
genetic applications, biochemical applications, pharmaceutical
applications, biomedical applications, medical applications,
radiological applications and environmental applications.
For medical applications, the sensing system and method of the
present embodiments is suitable for monitoring presence, and more
preferably level, of a biomarker in a biological liquid, such as a
physiological solution.
For environmental applications the sensing system and method of the
present embodiments is suitable for monitoring presence, and more
preferably level, markers indicative of the presence or level of
hazardous materials in a liquid, such as, but not limited to, water
pollutants, chemical agents, biological organisms or radiological
conditions in water.
The liquid can be a liquid that comprise blood product, either
whole blood or blood component. For example, the liquid can be a
blood sample. The liquid can comprise other body liquids,
including, without limitation, saliva, cerebral spinal fluid, urine
and the like. The liquid can be a buffer or a solution, such as,
but not limited to, nucleic acid solutions, protein solutions,
peptide solutions, antibody solutions and the like. Also
contemplated are liquids containing one or more biological and
chemical reagents such as, but not limited to, oxidizing agents,
reducing agents, enzymes, receptor ligands, extracellular
components, metabolites, fatty acids, steroids, and the like. A
representative list of liquids in which the system and method of
the present embodiments can sense a marker, include, without
limitation, water, salt water, urine, blood, sperm, saliva, mucous,
catemenial fluid, lymphatic fluid, cerebral spinal fluid, vaginal
exudate, pus, vomit, perspiration, and inorganic liquids,
including, without limitation, petroleum liquids, oils or other
lubricants.
According to some embodiments of the invention, a liquid, e.g.,
biological liquid having a marker, e.g., a biomarker, is introduced
into a fluidic system that comprises a sensor having an immobilized
affinity moiety interacting with the marker and being configured to
generate a detectable signal responsively to the interaction. The
interaction is optionally and preferably being characterized by a
K.sub.D which is equal or less than 10.sup.-5 M or a K.sub.D which
is equal or less than 10.sup.-6 M or a K.sub.D which is equal or
less than 10.sup.-7 M or a K.sub.D which is equal or less than
10.sup.-8 M or a K.sub.D which is equal or less than 10.sup.-9 M or
a K.sub.D which is equal or less than 10.sup.-10 M. The marker from
the liquid binds selectively to the affinity moiety. Background
components, which may include objects of any type, such as, but not
limited to, salts and bio-molecules, may be adsorbed
non-specifically to the sensor's surface.
After the selective adsorption of the marker to the sensor's
surface, the sensor is washed with a washing buffer. This results
in a desorption of background components from the surface of the
sensor. Since the interactions between the marker and the affinity
moiety are stronger than the interaction of the background
components with the surface of the sensor, the background
components leave the sensor's surface much faster than the marker.
The desorption kinetic of the marker from the sensor are preferably
detected after all or most of the background components are washed
out.
While the embodiments below are described with a particular
emphasis to biomarkers in a biological liquid, it is to be
understood that other types of markers and other types liquid are
also contemplated.
The system of the present embodiments optionally and preferably
provides a direct analysis of bio-samples on a single chip. The
system of the present embodiments can selectively detect specific
low abundant biomarkers, while removing unwanted components (salts,
bio-molecules, proteins, cells, etc.). Preferably, the analysis is
performed without performing at least one of, or more preferably
without any of: centrifugation, desalting and affinity columns,
since such operations are known to be time-consuming. In some
embodiments of the present invention the analysis process is
performed in less than 15 minutes or less than 10 minutes or less
than 5 minutes.
In some embodiments of the present invention the amount of the
biomarker in the bio-sample is also measured. This can be done, for
example, by providing a system in which some sensors include the
affinity moiety and some do not include the affinity moiety. The
desorption from the sensors that do not include the affinity moiety
is defined as a background. The desorption kinetic of the biomarker
from sensors that include the affinity moiety are optionally and
preferably compared to this background, and the amount of the
biomarker desorbed from sensors that include the affinity moiety is
determined based on this comparison, for example, by subtracting
the signals from each other. Since the specific absorption of the
biomarkers to the surfaces is proportional to the concentration of
the biomarker in the bio-sample, the desorption kinetics above
background is concentration-dependent.
Reference is now made to FIG. 7 which is a flowchart diagram
schematically illustrating of a method suitable for detecting a
presence and/or concentration of a marker, e.g., a biomarker, in a
liquid, e.g., a biological liquid, according to some embodiments of
the present invention.
It is to be understood that, unless otherwise defined, the
operations described hereinbelow can be executed either
contemporaneously or sequentially in many combinations or orders of
execution. Specifically, the ordering of the flowchart diagrams is
not to be considered as limiting. For example, two or more
operations, appearing in the following description or in the
flowchart diagrams in a particular order, can be executed in a
different order (e.g., a reverse order) or substantially
contemporaneously. Additionally, several operations described below
are optional and may not be executed.
The method begins at 10 and optionally and preferably continues to
11 at which the liquid is contacted with a sensor. The contact can
be established by any technique known in the art. Preferably, but
not necessarily, the sensor is in a fluidic system, more preferably
a microfluidic system, and the liquid is introduced into the
fluidic system.
FIGS. 8A and 8B are schematic illustrations of sensors suitable for
the present embodiments. FIG. 8A shows a sensor 20 which comprises
an immobilized affinity moiety 48. The sensor used in the technique
of the present embodiments can be any potentiometric sensor that
provides a detectable signal in the present of the marker or
biomarker.
The detectable signal is typically produced when the electrical
property (conductivity, resistivity, capacitance) of the sensor
varies in response to interaction with a marker or biomarker 50. In
some embodiments, the sensor is a transistor. In these embodiments,
moiety 48 is optionally and preferably immobilized on a surface of
the channel of the transistor.
According to some embodiments, the sensor is a nanostructure and
the affinity moiety is immobilized on a surface of the
nanostructure. According to some embodiments of the invention, the
sensor is a transistor, having a nanostructure as a channel,
wherein the affinity moiety is immobilized on a surface of the
nanostructure.
According to some embodiments, the sensor is a structure that is
non-nanometric. In these embodiments, all the dimensions of the
structure on which the affinity moiety is immobilized (length,
width and thickness) are above 1000 nm, or above 10 .mu.m, or above
100 .mu.m, or above 1 mm.
According to some embodiments of the invention, the sensor is a
transistor, having a non-nanometric structure as a channel, wherein
the affinity moiety is immobilized on a surface of the
nanostructure.
According to some embodiments of the invention, the transistor is a
field-effect transistor (FET).
When the sensor comprises a transistor (e.g., a FET) and the
affinity moiety is immobilized on the channel of the transistor, a
change in the electrical property of the channel can induce a
change in the characteristic response of the transistor to the gate
voltage (e.g., the source-drain current as a function of the gate
voltage), which change can be detected and analyzed.
FIG. 8A illustrates an embodiment in which sensor 20 comprises a
structure 40, such as, but not limited to, as nanostructure,
wherein the affinity moiety 48 is immobilized on a surface of
structure 40.
Affinity moiety 48 is effective to react (e.g., bind) specifically
to a marker or biomarker 50 in the liquid or biological liquid.
The sensor 20 is configured to generate a detectable signal,
typically an electrical signal, responsively to the interaction of
the affinity moiety 48 with the marker or biomarker 50.
As used herein the term "affinity moiety" refers to a molecule
which binds with a predetermined affinity and preferably
specificity to the marker or biomarker.
Affinity moiety 48 and marker or biomarker 50 are optionally and
preferably members of an affinity pair, wherein moiety 48 is
capable of reversibly or non-reversibly binding to marker or
biomarker 50. The interaction between moiety 48 and marker or
biomarker 50 is characterized by an affinity which is typically
weaker than the characteristic affinity of a covalent bond, which,
when expressed in K.sub.D, typically correspond to K.sub.D of about
10.sup.-15 M. In any of the embodiments of the invention, the
interaction between moiety 48 and marker or biomarker 50 is
characterized by an affinity which is preferably defined by a
K.sub.D that is equal or less than 10.sup.-5, or a K.sub.D that is
equal or less than 10.sup.-6M, or a K.sub.D that is equal or less
than 10.sup.-7M, or a K.sub.D that is equal or less than
10.sup.-8M, or a K.sub.D that is equal or less than 10.sup.-9M, or
a K.sub.D that is equal or less than 10.sup.-10 M.
Methods of measuring the affinity are well known in the art and
include surface Plasmon resonance and competition assays.
The affinity moiety may be naturally occurring or synthetically
designed or produced.
Examples of affinity moieties include a member of an
antibody-antigen (immunogenic moiety), a ligand-receptor (e.g.,
soluble receptor or membrane bound), a carbohydrate-lectin, an
RNA-aptamer, a nucleic acid sequence complementation and the
like.
According to some embodiments of the invention, the affinity moiety
comprises an immunogenic moiety. According to some embodiments of
the invention, the immunogenic moiety comprises an antibody or a
fragment thereof. According to some embodiments of the invention,
the immunogenic moiety comprises an antigen. In these embodiments,
the marker is a biomarker that preferably comprises an antibody to
the antigen. According to some embodiments of the invention, the
affinity moiety comprises a ligand. In these embodiments, the
marker is preferably a biomarker that comprises a receptor.
Moiety 48 can attached to the surface of nanostructure 40 by any
technique known in the art, such as, but not limited to, the
technique that is based on fragmentation of antibody-capturing
units and that is described in Elnathan et al., Nano Lett 2012, 12,
(10), 5245-5254, the contents of which are hereby incorporated by
reference.
Moiety 48 can be attached to the surface of nanostructure 40 by
means of reactive groups within moiety 48 and compatible reactive
groups on the surface of nanostructure 40, directly or via a
linker. Preferably the attachment is a covalent attachment. In
exemplary embodiments, the linker generates a reactive amine group
on the surface of the nanostructure, which is optionally subjected
to reductive amination to provide an aldehyde-terminated surface
that binds with moiety 48. The reactive groups on the surface of
nanostructure 40 can be intrinsic or can be generated upon a
treatment.
Nanostructure 40 is preferably elongated.
As used herein, a "elongated nanostructure" generally refers to a
three-dimensional body which is made of a solid substance, and
which, at any point along its length, has at least one
cross-sectional dimension and, in some embodiments, two orthogonal
cross-sectional dimensions less than 1 micron, or less than 500
nanometers, or less than 200 nanometers, or less than 150
nanometers, or less than 100 nanometers, or even less than 70, less
than 50 nanometers, less than 20 nanometers, less than 10
nanometers, or less than 5 nanometers. In some embodiments, the
cross-sectional dimension can be less than 2 nanometers or 1
nanometer.
In some embodiments, the nanostructure has at least one
cross-sectional dimension ranging from 0.5 nanometers to 200
nanometers, or from 1 nm to 100 nm, or from 1 nm to 50 nm.
The length of a nano structure expresses its elongation extent
generally perpendicularly to its cross-section. According to some
embodiments of the present invention the length of the
nanostructure ranges from 10 nm to 50 microns.
The cross-section of the elongated nanostructure may have any
arbitrary shape, including, but not limited to, circular, square,
rectangular, elliptical and tubular. Regular and irregular shapes
are included.
In various exemplary embodiments of the invention the nanostructure
is a non-hollow structure, referred to herein as "nanowire".
A "wire" refers to any material having conductivity, namely having
an ability to pass charge through itself.
In experiments performed according to some embodiments of the
present invention silicon nanowires, about 20 nm in diameter and
about 10 .mu.m in length, have been employed.
In some embodiments, a nanowire has an average diameter that ranges
from 0.5 nanometers to 200 nanometers, or from 1 nm to 100 nm, or
from 1 nm to 50 nm.
In some embodiments of the present invention, the nanostructure is
shaped as hollow tubes, preferably entirely hollow along their
longitudinal axis, referred to herein as "nanotube" or as
"nanotubular structure".
The nanotubes can be single-walled nanotubes, multi-walled
nanotubes or a combination thereof.
In some embodiments, an average inner diameter of a nanotube ranges
from 0.5 nanometers to 200 nanometers, or from 1 nm to 100 nm, or
from 1 nm to 50 nm.
In case of multi-walled nanotubes, in some embodiments, an
interwall distance can range from 0.5 nanometers to 200 nanometers,
or from 1 nm to 100 nm, or from 1 nm to 50 nm.
It is appreciated that while FIG. 8A shows a single nanostructure
40, some embodiments contemplate a configuration in which sensor 20
comprises a plurality (i.e., two or more) of nanostructure. When a
plurality of nanostructures is employed, the nanostructures 40 are
optionally and preferably arranged in an array. For example, the
nanostructures can be arranged generally parallel to each other, as
schematically illustrated in FIG. 8B.
Selection of suitable materials for forming nanostructure 40 as
described herein will be apparent and readily reproducible by those
of ordinary skill in the art, in view of the guidelines provided
herein for beneficially practicing embodiments of the invention.
For example, nanostructure 40 of the present embodiments can be
made of an elemental semiconductor of Group IV, and various
combinations of two or more elements from any of Groups II, III,
IV, V and VI of the periodic table of the elements.
As used herein, the term "Group" is given its usual definition as
understood by one of ordinary skill in the art. For instance, Group
III elements include B, Al, Ga, In and Tl; Group IV elements
include C, Si, Ge, Sn and Pb; Group V elements include N, P, As, Sb
and Bi; and Group VI elements include O, S, Se, Te and Po.
In some embodiments of the present invention the nanostructure is
made of a semiconductor material, optionally and preferably a
semiconductor material that is doped with donor atoms, known as
"dopant". The present embodiments contemplate doping to effect both
n-type (an excess of electrons than what completes a lattice
structure lattice structure) and p-type (a deficit of electrons
than what completes a lattice structure) doping. The extra
electrons in the n-type material or the holes (deficit of
electrons) left in the p-type material serve as negative and
positive charge carriers, respectively. Donor atoms suitable as
p-type dopants and as n-type dopants are known in the art.
For example, the nanostructure can be made from silicon doped with,
e.g., B (typically, but not necessarily Diborane), Ga or Al, to
provide a p-type semiconductor nanostructure, or with P (typically,
but not necessarily Phosphine), As or Sb or to provide an n-type
semiconductor nanostructure.
In experiments performed by the present inventors, Si nanowires and
p-type Si nanowires with a diborane dopant have been utilized.
In some embodiments of the present invention the nanostructure is
made of, or comprises, a conductive material, e.g., carbon. For
example, the nanostructure can be a carbon nanotube, either
single-walled nanotubes (SWNT), which are can be considered as long
wrapped graphene sheets, or multi walled nanotubes (MWNT) which can
be considered as a collection of concentric SWNTs with different
diameters. A typical diameter of a SWNT is less of the order of a
few nanometers and a typical diameter of a MWNT is of the order of
a few tens to several hundreds of nanometers.
When a plurality of nano structures is employed, the nanostructures
can be grown using, for example, chemical vapor deposition.
Alternatively, the nanostructures can be made using laser assisted
catalytic growth (LCG). Any method for forming a semiconductor
nanostructure and of constructing an array of a plurality of
nanostructures is contemplated. When a plurality of nanostructures
40 is employed, there is an affinity moiety 48 immobilized on each
of the nanostructures. In some embodiments of the present invention
all the affinity moieties are the same across all the
nanostructures, and in some embodiments at least two nanostructures
are attached to different affinity moieties.
A reaction event between marker or biomarker 50 and moiety 48
changes the surface potential of nanostructure 40 and therefore
results in a change of an electrical property of nanostructure 40.
For example, nanostructure 40 can exhibit a change in density of
electrons or holes over some region of nanostructure 40 or over the
entire length of nanostructure 40. Nanostructure 40 can
additionally or alternatively exhibit a change in its conductivity
or resistivity.
Referring again to FIG. 7, the method optionally and preferably
proceeds to 12 at which the liquid or biological liquid is washed
off the sensor, and to 13 at which the presence and/or
concentration of the marker or biomarker is detected based on a
detectable signal received from the sensor within a time-window
beginning a predetermined time period (e.g., at least 10 seconds or
at least 20 seconds or at least 30 seconds or at least 45 seconds
or at least 60 seconds or at least 75 seconds or at least 90
seconds or at least 105 seconds or at least 120 seconds or at least
135 seconds or at least 150 seconds) after the beginning time of
the washing 12. Preferably, the detection is based on signal
received within the time-window, but is not based on signal
received from the sensor before the beginning time of the
time-window. The duration of the time-window is preferably from
about 30 seconds to about 500 seconds. Other predetermined time
periods and time-window durations, including predetermined time
periods and time-window durations that are outside the above
ranges, are also contemplated.
According to some embodiments of the invention the signal is
monitored from the beginning of the washing, more preferably from
immediately before or immediately after the initiation of the
washing, but the beginning of the time-window during which the
signals on which the determination of the presence or level of the
marker is based, is not at the beginning of the washing. In these
embodiments, the method optionally and preferably determines the
beginning of the time-window from the signal itself. This can be
done, for example, by monitoring the time-dependence of the signal
(e.g., slope, plateau, zeroing of some derivative with respect to
the time, value of some derivative with respect to the time, etc.),
and identifying the beginning of the time-window based on a change
in the time-dependence. For example, the method can identify the
beginning of the time-window as a time point at which the signal
exhibits a decrement, or a time point at which the signal exits a
plateau region.
The detection is optionally and preferably by monitoring the change
in the electrical property of nanostructure 40 using an arrangement
of electrodes. With reference to FIGS. 8A and 8B, in some
embodiments of the present invention sensor 20 comprises a source
electrode 42 and a drain electrode 44, wherein nanostructure 40 is
disposed between electrodes 42 and 44 and serves as a charge
carrier channel. Optionally, sensor 20 also comprises a gate
electrode 46, forming, together with electrodes 42 and 44 and
nanostructure 40, a transistor, e.g., a field effect transistor
(FET). The gate electrode 46 is optionally and preferably, but not
necessarily, spaced apart from nanostructure 40 by a gap 47. A gate
voltage can be applied to channel nanostructure 40 through gate
electrode 46. In some embodiments, when the voltage of gate
electrode 46 is zero, nanostructure 40 does not contain any free
charge carriers and is essentially an insulator. As the gate
voltage is increased, the electric field caused attracts electrons
(or more generally, charge carriers) from source electrode 42 and
drain electrode 44, and nanostructure 40 becomes conducting. In
some embodiments, no gate voltage is applied and the change in the
charge carrier density is effected solely by virtue of the
interaction between affinity moiety 48 and marker or biomarker
50.
The electrodes of sensor 20 can be connected directly or indirectly
to a circuit (not shown). The circuit can apply voltage to
nanostructure 40 via one or more of the electrodes, and monitors
the changes in the electrical property of nanostructure 40
responsively to the binding of marker or biomarker 50 to affinity
moiety 48. The circuit can be constructed, for example, for
measuring an electrical measure corresponding to a change in the
electrical property of nanostructure(s) 40. The electrical measure
can be, e.g., voltage, current, conductivity, resistance,
impedance, inductance, charge, etc.
The method optionally and preferably continues to 14 at which the
liquid or biological liquid is contacted with another sensor, which
is preferably non-specific to the marker or biomarker 50. In these
embodiments, the method optionally and preferably continues to 15
at which the liquid or biological liquid is washed also off the
non-specific sensor. When operations 14 and 15 are executed, the
detection of the presence and/or concentration of the marker or
biomarker is optionally and preferably based on a comparison 16
between the detectable signal and a background signal received from
the non-specific sensor. A detection based on the comparison is
generally shown at 17.
The method ends at 18.
FIG. 9 is a schematic illustration of a system 90 for detecting a
presence and/or concentration of a marker or biomarker 50 (not
shown) in a liquid or biological liquid 92, according to some
embodiments of the present invention. System 90 can be used for
executing one or more, more preferably all, the operations of
method 10. System 90 can comprise a fluidic device 94 having a
sensing chamber 96 and a sensor, such as, but not limited to,
sensor 20, in the sensing chamber 96. System 90 can optionally and
preferably comprise a flow control system 98 for introducing a
washing buffer 100 to chamber 96 to wash liquid or biological
liquid 92 off sensor 20, and a signal analyzer 102 having a circuit
104 for analyzing detectable signals received from sensor 20, as
further detailed hereinabove.
In some embodiments of the present invention fluidic device 94
comprises a non-specific sensor 106, which is non-specific the
marker or biomarker, as further detailed hereinabove. Sensor 106
can be in the same sensing chamber 96 as sensor 20. In these
embodiments, the circuit 102 of signal analyzer 104 can compare
between the detectable signal and the background signal, and detect
the presence and/or concentration of the marker or biomarker based
on the comparison, as further detailed hereinabove.
As used herein the term "about" refers to .+-.10% or .+-.5%.
The terms "comprises", "comprising", "includes", "including",
"having" and their conjugates mean "including but not limited
to".
The term "consisting of" means "including and limited to".
The term "consisting essentially of" means that the composition,
method or structure may include additional ingredients, steps
and/or parts, but only if the additional ingredients, steps and/or
parts do not materially alter the basic and novel characteristics
of the claimed composition, method or structure.
As used herein, the singular form "a", "an" and "the" include
plural references unless the context clearly dictates otherwise.
For example, the term "a compound" or "at least one compound" may
include a plurality of compounds, including mixtures thereof.
The following presents some of advantages of the methodology of the
present embodiments over other technologies: Sensitivity from
femtoMolar to nanoMolar range (biomarker concentration in the
sample); Fast results, after less than about 5 minutes; Real-time
ultra-sensitive monitoring; Rapid and lack of time consuming
processes (i.e. centrifugation, dialysis, affinity columns);
Label-free: less steps to worry about, do not need to excite and
image; Multiplex: extract as much biological data as possible;
Overcoming the Debye screening length; Easy to integrate with a lab
on chip system; Work with very small volume sample;
Reusable/reversible; Low-cost.
Throughout this application, various embodiments of this invention
may be presented in a range format. It should be understood that
the description in range format is merely for convenience and
brevity and should not be construed as an inflexible limitation on
the scope of the invention. Accordingly, the description of a range
should be considered to have specifically disclosed all the
possible subranges as well as individual numerical values within
that range. For example, description of a range such as from 1 to 6
should be considered to have specifically disclosed subranges such
as from 1 to 3, from 1 to 4, from 1 to 5, from 2 to 4, from 2 to 6,
from 3 to 6 etc., as well as individual numbers within that range,
for example, 1, 2, 3, 4, 5, and 6. This applies regardless of the
breadth of the range.
Whenever a numerical range is indicated herein, it is meant to
include any cited numeral (fractional or integral) within the
indicated range. The phrases "ranging/ranges between" a first
indicate number and a second indicate number and "ranging/ranges
from" a first indicate number "to" a second indicate number are
used herein interchangeably and are meant to include the first and
second indicated numbers and all the fractional and integral
numerals therebetween.
Unless otherwise defined, all technical and/or scientific terms
used herein have the same meaning as commonly understood by one of
ordinary skill in the art to which the invention pertains. Although
methods and materials similar or equivalent to those described
herein can be used in the practice or testing of embodiments of the
invention, exemplary methods and/or materials are described below.
In case of conflict, the patent specification, including
definitions, will control. In addition, the materials, methods, and
examples are illustrative only and are not intended to be
necessarily limiting.
Implementation of the method and/or system of embodiments of the
invention can involve performing or completing selected tasks
manually, automatically, or a combination thereof. Moreover,
according to actual instrumentation and equipment of embodiments of
the method and/or system of the invention, several selected tasks
could be implemented by hardware, by software or by firmware or by
a combination thereof using an operating system.
For example, hardware for performing selected tasks according to
embodiments of the invention could be implemented as a chip or a
circuit. As software, selected tasks according to embodiments of
the invention could be implemented as a plurality of software
instructions being executed by a computer using any suitable
operating system. In an exemplary embodiment of the invention, one
or more tasks according to exemplary embodiments of method and/or
system as described herein are performed by a data processor, such
as a computing platform for executing a plurality of instructions.
Optionally, the data processor includes a volatile memory for
storing instructions and/or data and/or a non-volatile storage, for
example, a magnetic hard-disk and/or removable media, for storing
instructions and/or data.
Optionally, a network connection is provided as well. A display
and/or a user input device such as a keyboard or mouse are
optionally provided as well.
It is appreciated that certain features of the invention, which
are, for clarity, described in the context of separate embodiments,
may also be provided in combination in a single embodiment.
Conversely, various features of the invention, which are, for
brevity, described in the context of a single embodiment, may also
be provided separately or in any suitable subcombination or as
suitable in any other described embodiment of the invention.
Certain features described in the context of various embodiments
are not to be considered essential features of those embodiments,
unless the embodiment is inoperative without those elements.
Various embodiments and aspects of the present invention as
delineated hereinabove and as claimed in the claims section below
find experimental support in the following examples.
EXAMPLES
Reference is now made to the following examples, which together
with the above descriptions illustrate some embodiments of the
invention in a non limiting fashion.
Example 1
Fabrication of SiNW-FET
FIGS. 1A-G are images (FIGS. 1A and 1B) and schematic illustrations
of a fabrication process (FIGS. 1C-G) of a 20 nm diameter P-type
SiNW-FET device on 3 inch silicon wafer with 600 nm oxide layer,
according to some embodiments of the present invention.
P-type SiNWs were synthesized by chemical vapor deposition (CVD)
system (via vapor liquid solid (VLS) process) (FIG. 1C). The p-type
SiNWs were deposited on silicon substrate with 600 nm oxide layer
and outer metal pads (5 nm Cr and then 60 nm Au), that were
fabricated in advance by lithography (FIG. 1D). Source and drain
electrodes were deposited with the use of a multilayer photoresist
structure consisting of 500 nm LOR5A (Microchem) and 500 nm 1805
(Shipley). After exposure and development of the electrode patterns
(FIG. 1E), the contacts were metallized by e-beam and thermal
evaporation of Ni (60 nm) respectively, and were then passivated
with an insulating layer of Si.sub.3N.sub.4 (60 nm thick) deposited
by plasma-enhanced chemical vapor deposition at 80.degree. C.
(ICP-PECVD, Axic Inc.) and a layer of 10 nm alumina (ALD deposition
using a Cambridge Nanotech Savannah 200 system) (FIG. 1F). A
lift-off of un-exposed photoresists layers and a thermal annealing
of SiNWs and metal contacts were preformed (FIG. 1G). The
separation between the source and drain electrodes for each FET was
about 2 .mu.m. An image and a magnified image of the FET are shown
in FIGS. 1B and 1A, respectively.
Example 2
Preparation of SiNW Modified with Antibodies
A modification of silicon nanowires (SiNW) FET device with
antibodies is schematically illustrated in FIG. 2. The process is
briefly described as follows:
Surface Cleaning and Activation:
a SiNW FET prepared as described in Example 1 herein was washed
with acetone, deionized water and isopropanol; Dried under N.sub.2
stream; and the surface was thereafter activated for silanization
with oxygen plasma (0.200 torr, 100 W, 10 minutes).
Silanization:
The device surface was covered with
(3-aminopropyl)-dimethyl-ethoxysilane and heated to 50.degree. C.
for 60 minutes. Thereafter the surface was washed with isopropanol,
followed by dehydration on hot plate (115.degree. C., 30
minutes).
Cross Linker Binding:
The device with was covered with 8.3% glutaraldehyde containing 12
mM sodium cyanoborohydride in phosphate buffer 10 mM, pH=8.5 (60
minutes, room temperature). Thereafter, the device was washed with
deionized water, acetone, isopropanol and deionized water
again.
Channel Assembly on the Chip:
The tubing and PDMS channel were washed with isopropanol and
deionized water. The tubing was thereafter connected to a syringe
pump, and the system was washed by introducing phosphate buffer
(PB) (10 mM, pH=8.5).
Antibody Immobilization:
CA 15-3 igG (40 .mu.l, 1 mg/ml) was mixed gently with 700 .mu.l
phosphate buffer (10 mM, pH=8.5) containing 12 mM sodium
cyanoborohydride. The Antibody solution was introduced into the
system (at 4.degree. C., overnight about 16 hours), and the tubing
system was thereafter washed with phosphate buffer (10 mM, pH=8.5)
while keeping the surface under the channel always wet.
Blocking:
A blocking solution containing ethanolamine (100 mM) and 12 mM
sodium cyanoborohydride in PB (pH=8.5) was introduced into the
system during 3 hours at room temperature, at a flow rate of 50
.mu.l/minute. The system was thereafter washed with phosphate
buffer (10 mM, pH=8.5) at a flow rate of 50 .mu.l/minute, for 30
minutes. The as prepared system was used for sensing.
Example 3
Sensing
A fluidic system containing the a plurality of SiNW FET as
described in Example 2 above, was studied experimentally, using an
antigen-containing buffer, a washing buffer and a bio-sample. In
this example, the SiNW FET was modified with anti-troponin T (Fab2
fragment).
The antigen-containing buffer included the antigen cardiac troponin
T, 10 nM. The washing buffer included 150 .mu.M phosphate
buffer.
FIG. 3 presents an exemplary response calibration curve the
antibody-modified SiNW FET, upon introducing of an
antigen-containing buffer, following by introducing a washing
buffer into a fluidic system containing the a plurality of SiNW
FET.
The arrow A in FIG. 3 denotes the time point at which the
antigen-containing buffer was introduced into the fluidic system.
At this point, the antigen was absorbed on the SiNW surfaces due to
interactions with the antibody, and a change in the electric
response is detected. The arrow B in FIG. 3 denotes the time point
at which the washing buffer was introduced into the fluidic system.
At this point, most of the bound antigen was desorbed from the
SiNWs, and a reverse change in electric response is detected. The
parameters used for the FET are Vg=0 volt, Vds=0.1 volt. The flow
rate in the fluidic system was 20 .mu.l/min, generated by Mitos
pump.
FIG. 4 presents calibrated response of antibody-modified SiNW FET,
in a bio-sample. In this experiment, the bio-sample was fetal
bovine serum. The arrow A in FIG. 4, denotes the time point at
which the bio-sample was introduced into the system. The parameters
used for the FET are Vg=0 volts, Vds=0.2 volts. The flow rate in
the fluidic system was 100 .mu.l/min.
FIG. 5 presents a response calibration curve in the presence of the
bio-sample of FIG. 4, but after initiation of a washing operation.
The arrow B in FIG. 5, denotes the time point at which the washing
buffer was introduced into the system. The parameters used for the
FET and the flow rate were the same as in FIG. 4 above. A
concentration of less than 30 units/ml CA 15-3 produces a
detectable change in electric response.
FIG. 6 shows two response calibration curves obtainable in the
presence of a bio-sample, after the initiation of a washing
operation. The arrow B in FIG. 6 denotes the time point at which
the washing buffer was introduced into the system. A first response
calibration curve corresponds to the desorption of the biomarker
from a sensor with affinity moiety, and the second calibration
curve corresponds to the desorption of the biomarker from a sensor
without affinity moiety. The second calibration curve can be
defined as the background to which the first curve can be
compared.
Example 4
Detailed Study
This example demonstrate the application of antigen-dissociation
regime, from antibody-modified Si-nanowire sensors, as a simple and
effective direct sensing mechanism of biomarkers of interest in
complex biosamples, such as serum and untreated blood, which does
not require ex situ time-consuming biosample manipulation steps,
such as centrifugation, filtering, preconcentration, and desalting,
thus overcoming the detrimental Debye screening limitation of
nanowire-based biosensors.
A fluid-delivery device was fabricated from flexible
polydimethylsiloxane (PDMS) elastomer. The PDMS was incubated with
curing agent at 10:1 mass ratio for overnight at 60.degree. C. The
resulting device was then cut into rectangular pieces, at
dimensions of 10.times.10.times.5 mm. Two channels of different
dimensions were used in this study: a rectangular chamber of larger
dimensions (h=5 mm, 1=7 mm and w=3 mm), and a rectangular smaller
chamber (h=0.1 mm, 1=3.5 mm and w=1 mm). An upstream polyethylene
tube (PE 20, Intramedic) was 14 cm long and had 0.38 mm inner
diameter. A downstream Tygon tube (S-50-HL, Tygon) was 13 cm
long.
A chip with an array of SiNW FET were chemically-modified to
perform sensing of binding and unbinding kinetics of antigen by
immobilized antibody on the SiNW FET surface. In order to conjugate
the antibody to the SiNWs surface, the chip was first washed with
acetone (9005-68, J. T. Baker), isopropanol (9079-05, J. T. Baker),
and deionized water (18 M.OMEGA.cm) successively, followed by
nitrogen drying. Then, oxygen plasma (100 W, 0.2 Torr) was applied
for 10 minutes. The chip was covered by glass dish and inserted to
glove box (150B-G, Mbraun) under argon atmosphere (water and oxygen
free) to apply the amino-silane modification. Immediately
afterwards, the chip was covered with about 150 .mu.l
(3-aminopropyl)-dimethyl-ethoxysilane (APDMES; SIA0603.0, Gelest)
for 60 minutes. Then, the chip was washed three times with about 30
ml of anhydrous toluene (99.8%, 244511, Sigma-Aldrich). The chip
was transferred from the glove box to the clean room, and washed
again with isopropanol followed by nitrogen drying. Next, the chip
was placed on a hot plate at 115.degree. C. for 30 minutes. The
subsequent cross-linker binding was performed by covering the
device with 8.3% glutaraldehyde solution, containing 12 mM sodium
cyanoborohydride in 10 mM phosphate buffer, pH=8.5, for 60 minutes
at room temperature, followed by subsequent washes with deionized
water, acetone, isopropanol and deionized water.
PDMS channels that were pre-washed with isopropanol and deionized
water, were then assembled to the chip by connecting the tubing to
the syringe pump, followed by withdrawing 10 mM phosphate buffer,
pH=8.5. The antibody was immobilized to the SiNW surface by
withdrawing the antibody solution, containing 10-100 .mu.g/ml IgG
antibody, 12 mM sodium cyanoborohydride and 10 mM phosphate buffer,
pH=8.5, into the system at 4.degree. C. for overnight (about 16
hours). Blocking was performed by withdrawing the blocking
solution, containing 100 mM ethanolamine and 12 mM sodium
cyanoborohydride in 10 mM phosphate buffer, pH=8.5, into the system
at a flow rate of 50 .mu.l/min for 150 minutes at room temperature,
followed by final wash with 10 mM phosphate buffer, pH=8.5, at a
flow rate of 50 .mu.l/min for 30.
The devices were wire-bonded (using wire-bonder, model 8850, West
Bond) and the sensor device chip was integrated with the
custom-made PDMS microfluidic channel. A data acquisition system
was used to measure the current of the SiNW FETs (Ids), induced by
surface charges alterations. The selected devices were examined for
their performance in sensing buffer. Gate voltage sweep was used
for transconductance measurements, and the subsequent determination
of the transistor regime of operation. A suitable gate voltage was
further selected to perform all the following sensing experiments.
Sensing experiments were performed by monitoring the conductance of
the SiNW devices over time (current-versus-time signals were
recorded at 1 second intervals), during introduction of the
analytes to the sensing chip by a syringe pump (Fusion 200, Chemyx)
via the microfluidic system.
FIGS. 22A and 22B show electrical characterization of p-type SiNW
FET nanodevices under water-gate configuration. FIG. 22A is a plot
of source-drain current versus source-drain voltage (Vsd) at
different gate voltages (Vg). FIG. 22B is a plot of source-drain
current versus gate voltages (Vg) at 0.1 V source-drain voltage
(Vg).
FIGS. 23A and 23B are Scanning Electron Microscope (SEM) images of
the SiNWs. FIG. 23A is SEM image of the synthesized 20 nm p-type
SiNW via chemical vapor deposition system on silicon (100) wafer,
and FIG. 23B is a SEM image of SiNW FET device consisting of SiNWs
connected to source and drain electrodes.
It was found that the parameters that control the capability to
perform quantitative biomarkers analysis in biosamples include (i)
the affinity strength (k.sub.off rate) of the antibody-antigen
recognition pair, which dictates the time length of the
high-affinity slow dissociation sub-regime, and (ii) the flow rate
applied during the solution exchange dissociation step, which
controls the time width of the low-affinity fast dissociation
sub-regime. The lack of ex situ biosample manipulation
time-consuming processes enhances the portability of the sensing
platform and reduces to minimum the required volume of tested
sample, as it allows the direct detection of untreated biosamples
(5-10 .mu.L blood or serum), while readily reducing the detection
cycle duration to less than 5 min, factors of great importance in
near-future point-of-care medical applications.
Analysis of the dissociation regime of an antigen from its specific
antibody-modified SiNW FET device is demonstrated as an effective
and straightforward approach for the sensitive, selective, and
direct detection of biomarkers from complex biosamples.
The biosample containing the analyte biomarker was introduced to
the SiNW FET sensing nanodevice. The surface of the SiNW modified
with the antibody specifically interacted with the analyte
biomarker. As a result of antibody-antigen interaction, the
biomarker molecules strongly and selectively bind to the antibody
units attached to the surface of the nanowire FET device. In
physiological solutions such as serum or blood, it may be difficult
to discern the specific binding of the biomarker analyte cannot be
directly from the background signal change, due to the charge
screening caused by the high concentration of charged chemical
species, such as salts and proteins.
After the selective association of the biomarker molecules, the
device is rapidly flushed out with a controlled solution of low
ionic strength, the sensing buffer, wherein the SiNW FET device can
effectively sense the change in the surface charge on the nanowire,
due to the strongly bound antigen species, without the masking
effects of unbound nonspecific chemical species. Rapid washing of
the sensing devices with the sensing buffer results in the fast
removal of unbound, or loosely bound, nonspecific chemical species
(salts, proteins, cells, small molecules) from the nanowire surface
proximity, leaving behind only specific antigen molecules attached
via strong and specific interactions to the nanowire surface,
revealing considerably slower dissociation kinetics.
This effect efficiently splits the "dissociation regime window"
into two sub-regimes: (i) At the beginning, when the low ionic
strength "sensing buffer" is flushed through to the SiNWs FET
sensing array, the low-affinity entities (salts, biomolecules,
proteins, etc.) speedily leave the SiNWs surface, which results in
a rapid change in the conductivity of the devices. (ii) After the
removal of the low affinity entities ends (the unspecific
dissociation sub-regime), the specific dissociation of
high-affinity entities (the specific desorption sub-regime)
dominates the change of the conductivity of the device. The point
of transition between the unspecific and specific dissociation
sub-regimes is finally applied for the sensitive and accurate
detection of biomarker proteins.
This approach represents the direct analysis of complex biosamples
on a single platform, able to selectively detect low-concentration
specific biomarkers, while easily removing unwanted chemical
species (salts, biomolecules, proteins, cells), without the
requirement for time-consuming steps such as centrifugation,
desalting, or affinity columns. The whole ultrasensitive protein
label-free analysis process can be practically performed quickly,
for example, in less than 5 min.
FIG. 10 schematically describes the operation principle of the
dissociation regime detection approach of an antigen from
antibody-modified SiNW FET sensing devices, according to some
embodiments of the present invention. Initially, the immobilized
SiNW device is introduced to low ionic strength sensing buffer
aiming to achieve a stable baseline. Next, a biosample is
introduced either to a SiNW FET device that is modified with a
specific antibody, or to a SiNW FET device that is modified with
nonimmuno active protein, FIG. 10, section (1).
The interaction of the SiNW FET device with biomolecules in the
analyzed sample alters the conductivity of the device during the
binding regime window, FIG. 10, section (2). At this point, the
SiNW FET device cannot distinguish between the change in the
conductivity that is caused by the binding of the specific antigen
or by other biomolecules and salts in the sample, due to charge
screening effects in high ionic strength solutions. When the SiNW
FET device is introduced back to low ionic strength sensing buffer,
through rapid flush out of the sensing chamber, the unbound
biomolecules and salts from the biosample are removed from the
surface of the SiNW FET device in a manner that depends upon the
strength of the interaction between them.
In the case of SiNW FET devices modified with an antibody, the
presence of the specific antigen will reduce the rate of returning
to baseline, due to the much slower desorption kinetics of the
antigen from its specific antibody, FIG. 10, section (3). However,
when the SiNW surface is modified with a nonspecific antibody, or
when there is no antigen in the sample, the rate of returning to
baseline is considerably faster, due to the absence of specific
high-affinity interactions.
FIG. 11 shows normalized association response of a representative
anti-CA 15-3 immobilized SiNW FET sensing device against various
concentrations of the target antigen CA 15-3 in unprocessed serum
samples. The black curve represents the response of the FET device
to CA 15-3-free fetal bovine serum control sample. The red, blue,
and turquoise curves represent the responses of the FET sensing
device to samples containing CA 15-3 antigen concentrations of 55,
135, and 535 pM in unprocessed fetal bovine serum, respectively.
The black arrow depicts the sensing device electrical baseline
under the flow of low ionic-strength sensing buffer (sensing
buffer-SB, 155 .mu.M sodium phosphate buffer pH abut 8.0) through
the microfluidic channel before the injection of biosamples. The
blue arrow indicates the time of injection of the unprocessed serum
biosamples (high ionic-strength samples). Non-analytical sensing
information can be extracted from the association curves under
these conditions. The Inset of FIG. 11 shows raw data curves of the
interaction of the same sensing device against biosamples
containing different concentrations of the CA 15-3 target
antigen.
The Debye length limitation predicts that the nano-FET device will
not be able to sense the antigen-antibody high-affinity
interactions under high ionic strength physiological conditions
(Debye length is approximately 1 nm). Thus, the measured antigen
association curves cannot be translated to analytical signals (see
FIG. 11). The association curves obtained in serum samples
demonstrate high similarity between the normalized electrical
responses of the monoclonal antibody-modified nanowire device
against various concentrations of the CA 15-3 antigen (human-cancer
associated antigen), from 0 to 535 pM, which is a biomarker for
breast cancer diagnosis and monitoring. No analytical differences
are observed for the interaction of the SiNW FET sensing devices
against different concentrations of the tested CA 15-3 antigen in
unprocessed serum samples, due to the high ionic strength of the
biosample, thus preventing the use of the association-regime curves
as an analytical means for the real time detection of biomolecular
species. These results, therefore, demonstrate the strong
requirement for the development of novel nano-FET-based detection
approaches for the direct detection of protein biomarkers in
untreated biosample solutions, e.g., blood.
FIG. 12 shows normalized electrical response of a representative
antitroponin antibody-modified SiNW FET sensing device to the
association and dissociation of its specific antigen troponin T
(cTnT, 1 nM) under low-ionic strength conditions (sensing buffer,
SB, 155 .mu.M sodium phosphate buffer pH of about 8.0) (red curve).
Also shown (Black curve) is a normalized electrical response of a
representative antitroponin antibody-modified SiNW FET sensing
device against a BSA containing sample (bovine serum albumin, 1 nM)
under low ionic strength conditions (sensing buffer, SB, 155 .mu.M
sodium phosphate buffer pH of about 8.0). The black arrow depicts
the sensing device electrical baseline under the flow of low
ionic-strength sensing buffer (sensing buffer-SB, 155 .mu.M sodium
phosphate buffer pH of about 8.0) through the microfluidic channel
before the injection of the corresponding samples. The blue arrow
indicates the time of introduction of either cardiac troponin T
(the antigen, marked by a red line) or bovine serum albumin (the
serum protein, marked by a black line) samples to the SiNW FET
devices (association step). The green arrow indicates the
subsequent washing step with sensing buffer (dissociation
step).
In the specific case demonstrated in this example, cardiac troponin
T, a very important marker of heart failure, was used as an
antigen, while the antigen-binding monoclonal antibody against
cardiac troponin T (cTnT) was used as the specific antibody
receptor immobilized to the SiNW FETs surface. Initially, the SiNW
FET devices are exposed to sensing buffer solution, SB, and the
devices' electrical responses are normalized according to EQ. 1,
below 100%.times.(I.sub.SB-I.sub.t)/I.sub.SB. (EQ. 1) where
I.sub.SB is the current in sensing buffer (SB) and It is the
current at a certain time point during the measurement.
Next, a sample containing cardiac troponin T antigen in sensing
buffer was introduced into the detection channel (indicated in FIG.
12 by blue arrow), leading to a sharp increase in the concentration
of free troponin molecules in the close vicinity of the
nanodevices' surface, followed by the gradual high-affinity
association of troponin molecules to the surface-immobilized
antibody units. As a result, the normalized electrical response of
the devices increases, until reaching a saturation point or
plateau, indicating the maximum amount of bound troponin antigen to
the nanodevices' surface. Next, the sensing nanodevices were
rapidly flushed with the low ionic-strength sensing buffer. The
amount of bound troponin molecules gradually decreased in response,
until reaching the original sensing buffer baseline electrical
signal. Noteworthy, the observed dissociation kinetics of troponin
antigen is considerably slower than its association kinetics. These
observations are consistent with the expected high-affinity
interactions between the immobilized antibody receptor units and
the antigen molecules.
When exposing the SiNW FET devices, modified with anti-cTnT
antibody receptor units, to a low ionic strength solution
containing a high concentration of the protein BSA (bovine serum
albumin), no significant change in the conductivity of the
respective devices was observed. These findings demonstrate that
the association of protein molecules to the antibody-modified
SiNW-based sensing devices is dominated by specific high affinity
antigen-antibody interactions, and that the surface chemistry on
the nanowires prevents the nonspecific binding of biomolecules
other than the specific target antigen molecules. The antigen
association curves obtained under low-ionic strength conditions
allow extracting the saturation time where the association of the
antigen molecules reaches a plateau. All further dissociation
experimental data was achieved after antigen association reaches a
plateau, for consistency purposes.
By performing a rapid flush-out step of the antibody-modified SiNW
devices using a low ionic strength sensing buffer, after complete
association of the specific antigen molecules has been reached, the
device is capable to analytically split the dissociation regime
time window into two discrete dissociation sub-regimes: (i) a
fast-dissociation kinetics sub-regime related to low-affinity
interacting chemical entities (e.g., salts, nonspecific proteins,
cells, and small chemical species), and (ii) a slow-dissociation
kinetics sub-regime related to high-affinity interacting chemical
entities (e.g., the specific antigens).
In other words, the flushing operation using the low ionic strength
sensing buffer allows for the simultaneous fast removal of
low-affinity fast-dissociating unbound, or loosely bound,
nonspecific molecules from the close vicinity of the nanowires
surface, thus minimizing the charge screening effects caused by
these charged species, accompanied by the resulting capability to
measure the presence of high-affinity slow dissociating bound
antigen species, after the complete removal of low-affinity species
is achieved. A considerably contrasting dissociating kinetics rate
for the fast-dissociating species, in comparison with the
dissociation rates of the slow dissociating species,
k.sub.off.sup.antigen>>k.sub.off.sup.nonspecific species,
allows to experimentally split the dissociation window regime into
the above-mentioned sub-regimes. Thus, both the intrinsic affinity
constant of the selected surface-attached antibody, as well as the
applied flow rate during the SB flushing step can be selected to
perform accurate quantitative detection of the antigen species
based on the dissociation regime window.
In various exemplary embodiments of the invention the rapid ionic
strength exchange step using the SB solution does not affect
significantly the amount of specifically bound antigen species.
This can be ensured, for example, by selecting the time window such
that the antigen-antibody pair dissociation time is considerably
longer than the time-window for complete removal of unbound, and
weakly bound, nonspecific species (the nonspecific dissociation
sub-regime). The nonspecific species dissociation time window can
be controlled and shortened by increasing the flow rate of the
flushing SB solution during the dissociation regime window.
The antigen dissociation regime window was tested as a potential
means for the quantitative detection of biomarkers in complex
biosamples, for two model antibody systems characterized by highly
dissimilar affinity/dissociation constants. First, an anti-troponin
antibody fragment, which represents a structurally modified
fragment of the whole wild-type antibody molecule with a lower
association affinity against the antigen than the original whole
antibody, was employed aimed at testing our hypothesis. The results
are shown in FIG. 13 which shows anti-troponin antibody
dissociation curve in serum, together with a fit, under slow-flow
conditions. The dissociation rate, off-rate k.sub.off, for the
antitroponing antibody fragment was extracted from the antigen
dissociation curves measured under low ionic strength SB, and
estimated to be k.sub.off=1.2.times.10.sup.-2 s.sup.-1.
FIGS. 14A and 14B show results of a comparison between the
dissociation kinetics of the troponin T antigen-containing serum
sample (at a troponin concentration of 10 nM) from the
(F(ab')2)-immobilized SiNW FET devices and control troponin T-free
serum sample. Measurements were conducted as follows. At first, the
sample (either troponin T-containing serum or troponin T free serum
control sample) was introduced to the immobilized SiNW FET devices
through a microfluidic channel, at a flow rate of 100 .mu.L/min (1
chamber volume exchange per minute, chamber volume 100 .mu.L) until
reaching the association plateau. Next, the immobilized SiNW FET
devices are washed with the low ionic strength sensing buffer. The
resulting dissociation kinetic curves demonstrate a similar
temporal change of the normalized electrical response for both, the
troponin T-containing serum sample and the troponin-free control
sample (FIGS. 14A and 14B, red and black curves). The nanoFET
devices show a relatively fast return to the baseline electrical
response (under low ionic strength sensing buffer) and similar
dissociation-related temporal electrical responses, implying the
concurrent dissociation of nonspecific species together with the
specific biomarker molecules, with an apparent k.sub.off of about
2.5.times.10.sup.-2 s.sup.-1 for the antigen-free serum sample.
FIG. 15 shows the serum dissociation curve, together with a fit,
under slow-flow conditions.
The low-affinity anti-troponin antibody receptor selected is thus
less suitable for effective splitting of the dissociation regime
window and the quantitative detection of the antigen molecules in
the complex biosample.
Next, an antibody receptor with a significantly stronger binding
affinity to its specific antigen was used. The use of a
higher-affinity antibody, characterized by a lower dissociation
rate from its antigen, leads longer dissociation times of the
biomarker protein from the antibody-immobilized SiNW devices, hence
enabling the separation of the low-affinity fast dissociation
sub-regime (ionic species, nonspecific proteins, small chemicals,
and cells) from the high-affinity slow dissociation sub-regime (the
antigen). In the present example, a mouse monoclonal antihuman
cancer antigen 15-3 IgG (Anti-CA 15-3), was used. FIG. 16 shows
anti-CA 15-3 antibody dissociation curve in serum, together with a
fit) under slow-flow conditions. The obtained k.sub.off was
k.sub.off=6.2.times.10.sup.-4 s.sup.-1 indicating strong binding
interaction against its antigen.
FIGS. 17A-D shows concentration-dependent sensing of the CA 15-3
antigen in unprocessed serum samples using the dissociation regime
mode approach, at a low flow rate of 1 chamber-volume exchange per
minute. In FIG. 17A, the black, red, blue, and turquoise curves
show the complete raw electrical response of a representative
anti-CA 15-3 modified SiNW FET sensing device to the association
and dissociation (raising phase) of its specific antigen CA 15-3 in
unprocessed bovine serum sample at concentrations of 0 (control
antigen-free sample), 55, 135, and 535 pM, respectively. In FIG.
17B, the black, red, blue and turquoise curves show the normalized
dissociation regime electrical response of the anti-CA
15-3-modified SiNW FET sensing device in unprocessed bovine serum
sample at CA 15-3 concentrations of 0 (control antigen-free
sample), 55, 135, and 535 pM, respectively. Each antigen sample was
flowed for about 6 min (until reaching association plateau) before
the washing-out with low-ionic strength sensing buffer was
performed (sensing buffer-SB, 155 .mu.M sodium phosphate buffer pH
of about 8.0). The black arrow indicates the time of solution
exchange from the unprocessed serum samples to the low-ionic
strength sensing buffer. FIG. 17C shows concentration-dependent
calibration curves of the CA 15-3 antigen extracted at different
points of time along the measured dissociation regime curves in
FIG. 17B. The black squares, red circles, and blue, green, pink and
brown triangles correspond to points in time depicted as black,
red, blue green, pink and brown dashed lines in FIG. 17B,
respectively. The curves show that the analytically relevant high
affinity regime, marked in turquoise circles, begins after 450 s
along the dissociation curves in FIG. 17B. The red circles depict
non-analytically relevant points in time along the dissociation
regime curves. FIG. 17D shows concentration-dependent calibration
plot of the CA 15-3 antigen, at the analytically relevant time
points marked by turquoise circles in FIG. 17C).
FIGS. 17A-D demonstrate that the selection of a higher affinity
antibody receptor allows for a deconvolution of the dissociation
regime window into the two dissociation sub-regimes, an unspecific
fast dissociation sub-regime and a specific slow dissociation
sub-regime, with an inter-regime transition time of approximately
250 s at a flow rate of 100 .mu.L/min (1 chamber volume exchange
per minute, chamber volume 100 .mu.L).
From the resulting curves (FIG. 17A), it is discernible that the
dissociation window consists of two clearly separated sub-regimes,
a fast dissociation zone lasting approximately 250 s, followed by a
slow dissociation plateau-like zone that lasts for tens of minutes
without a considerable change, and different in its amplitude from
the electrical baseline signal of the devices prior interaction
with the antigen under the low-ionic strength buffer. This
observation demonstrates that the antigen molecules stay bound to
the nanowire-immobilized antibody species for a period of time
considerably longer than the time frame required for complete
removal of the low-affinity nonspecific chemical species,
k.sub.off.sup.antigen>k.sub.off.sup.nonspecific species. This
latter dissociation zone, characterized by an electrical response
higher than the sensing buffer baseline electrical response (FIG.
17B, horizontal dashed blue line) displays a strong and
reproducible dependence on the concentration of the antigen protein
CA 15-3 in serum, which remains tightly bound to the SiNW FET
device surface.
A further experimental evidence on the presence of the protein
antigen confined to the surface of the nanodevices was demonstrated
by the use of a regeneration buffer (glycine buffer, pH=3), which
rapidly brings the dissociation of the high-affinity
antigen-antibody pairs and causes the electrical response of the
nanodevices to return to their initial baseline electrical level,
after the subsequent flow of the low ionicstrength sensing
solution. This is shown in FIG. 18, which is a graph of the
regeneration curve of a CA 15-3 antigen from its antibody-modified
nanowire device.
Additionally, concentration dependent experiments performed on
untreated serum samples spiked with different concentrations of the
CA 15-3 antigen reveal the robustness of this quantitative
detection approach based on the simple examination of the
dissociation regime window. As the concentration of CA 15-3 in the
tested serum sample increases (FIG. 17B), more of the biomarker
molecules associate to the SiNW devices surface, leading to stable
larger electrical response (in relation to the SB baseline
response).
The strong specific interaction between the anti-CA 15-3 and the
antigen CA 15-3 allows the complete washing of the nonspecifically
adsorbed salts and biomolecules within the sensing channel,
performed by the fast flushing with low ionic strength sensing
buffer, while maintaining the majority of surface-bounded CA 15-3
antigen molecules, in order to measure their sample concentration
quantitatively.
Therefore, monitoring the dissociation regime of antigens from
antibody-immobilized surfaces, with high-affinity capabilities,
allows performing direct analytical detection using SiNW FET-based
device, without applying any sample manipulation steps. This allows
performing quantitative protein measurements, sensitive enough for
the clinical diagnostics of CA 15-3 in relevant physiological
concentration range (greater than 67 pM).
Thus, antibody species displaying dissociation rates in the range
between k.sub.off of about 5.times.10.sup.-3 s.sup.-1 and about
1.times.10.sup.-8 s.sup.-1 can serve according to some embodiments
of the present invention as detection receptors.
Through a controlled increase of the dissociation flow rate step,
the nonspecific species dissociation sub-regime window can be
selected such that antibodies of lower affinity may also be
used.
In various exemplary embodiments of the invention complete
association of the antigen species, for example, as verified by an
electrical signal plateau, are achieved before the dissociation
begins. This is advantageous when concentration dependent
analytical sensing results is desired, since the final
concentration of the antigen under test is calculated based on the
difference between the baseline electrical signal under low
ionic-strength conditions, prior to the antigen association, and
the dissociation curve plateau achieved after complete removal of
the low-affinity species from the vicinity of the sensing devices
(high-affinity regime). Thus, changes in the amount of associated
antigen species to the nanosensing devices, due to differences in
the antigen-association time applied during the sensing may cause
changes to the extracted concentration-dependent calibration
curve.
Thus, in various exemplary embodiments of the invention a constant
antigen association period of time, or, alternatively a complete
association allowed, is applied for analytical consistency.
To quantitatively assess the antigen concentration, the transition
time between the dissociation sub-regimes was firstly measured by
exposing the sensing devices to an antigen-free serum sample. This
sample served as reference for the extraction of the accurate
time-point in which a complete removal of the low affinity species
was achieved. After this point in time, an accurate quantitative
assessment of the antigen concentration can be confidently
performed, as demonstrated in FIGS. 17C and 17D, under the
assumption of slow dissociation of the antigen species.
It was found by the inventors that such calibrating steps are not
necessary. This will now be explained with reference to FIGS. 19A
and 19B, which demonstrate multiplex single-chip differential
detection of the CA 15-3 antigen by the use of specific and
nonspecific chemically modified SiNWs FET devices, at a flow rate
of 330 chamber-volumes per minute. Thus, the sensing arrays of the
present embodiments can be fabricated with two main types of
chemically modified nanodevices, wherein the first group represents
the sensing nanodevices, chemically modified with the antibodies
specific against the antigens under examination, and wherein the
second group of nanodevices is chemically modified with a nonimmune
reactive protein (or a nonspecific antibody receptor) and serves as
on-chip internal reference devices.
The latter group of devices, due to the absence of specific
interactions with the antigens in the biosample under test, only
nonspecifically interact with low-affinity fast dissociating
species present in the biosample, and allows simple extraction of
the accurate transition time at which the first dissociation
sub-regime is reached and a quantitative assessment of antigen
concentration can be carried out. Thus, using these nonreactive
on-chip reference devices allows for the simultaneous sensitive and
quantitative detection of biomarkers in real time mode. In
addition, increasing the flow rate of the dissociation-related
flushing step can lead to a narrowing of the low affinity
fast-dissociation sub-regime time window, and thus allow for a
faster and more accurate quantitative detection of the antigen
species. For this purpose a microfluidic chamber of smaller
dimensions was used, so as to allow for considerably higher nominal
flow rates (chamber volume exchange rate), using flow rates easily
achievable with the fluidic pumping system.
Additionally, the use of smaller dimension microfluidic channels,
instead of the previously used 100 .mu.L larger chamber, can lead
to a more efficient fluid exchange during the dissociation washing
regime, along with the critical requirement of considerably lower
biosample volumes, possibly lower than a few microliters. FIGS. 19A
and 19B demonstrate the measurements performed aimed at the
detection of the CA 13-5 biomarker based on the differential
on-chip detection approach discussed above, this time using a
considerably higher effective flow rate of 100 .mu.L/min (chamber
volume exchange rate is 330 chamber volumes per minute, chamber
volume 0.3 .mu.L).
By comparing the dissociation curves obtained from the nonimmune
active protein-modified nanowire devices to the dissociation curves
attained by the specific antibody-modified devices, the amount of
the biomarker associated with the antibody-modified nanowire-based
devices can be measured. The SiNW device modified with the protein
BSA, which does not have specificity against the biomarker CA 15-3,
reaches a plateau about 25 s after the flushing of the sensing
buffer (FIG. 19A, red and blue curves). The antigen-free serum
dissociation curve under fast-flow conditions is shown, together
with a fit, in FIG. 20. The calculated k.sub.off was about
1.4.times.10.sup.-1 s.sup.-1.
The desorption kinetics from the antibody-modified SiNW device is
considerably slower and correlates well with the concentration of
the targeted antigen. The application of a faster solution exchange
rate narrows the fast-dissociation sub-regime window by a 10-fold
factor, from about 250 s to about 25 s, thus allowing for a
considerably faster detection cycle without the requirement of
off-line calibration steps, while not affecting the quantitative
and sensitive accurate assessment of the antigen biomarkers. These
results demonstrate that the simultaneous combination of
bioreceptors of suitable k.sub.off values, along with the use of
microfluidic chambers of adequate dimensions (that allow the
fastest possible fluid exchange), allows for the direct fast,
sensitive, and accurate detection of biomolecules based on their
dissociation regimes.
The faster flow condition allows using antibody receptors, or other
bioreceptors, of considerably lower affinity. The technique
described in this Example can be applied to untreated blood
samples. The measurements performed in this example demonstrate
that a complete removal of the low-affinity nonspecific species
(blood cells, proteins, salts, and small chemicals) can be achieved
after a period of about 300 s, under low flow conditions, and
lasting shorter, about 80 s under high flow conditions. FIG. 21
shows antigen-free untreated Blood dissociation curve under
slow-flow conditions, together with a fit. The obtained k.sub.off
was 1.6.times.10.sup.-2 s.sup.-1.
The present example demonstrates the application of
antigen-dissociation regime from antibody-modified nanowire sensors
for use in direct sensing in complex biosamples, serum and
untreated blood. The technique of the present embodiments does not
require ex situ time-consuming biosample manipulation steps, such
as filtering, preconcentration, and desalting.
The combination of high-affinity antibody receptors, along with
high solution-exchange flow rates, leads to the effective
deconvolution of the complex dissociation regime window into two
fully separated dissociation sub-regimes, thus allowing
quantitative detection of biomarkers. The lack of ex situ biosample
manipulation processes enhances the portability of the sensing
platform and reduces the required volume of tested sample as it
allows the direct detection of untreated biosamples (for example,
from about 5 to about 10 .mu.l .mu.L blood or serum), reducing the
detection cycle duration to less than 5 min.
Although the invention has been described in conjunction with
specific embodiments thereof, it is evident that many alternatives,
modifications and variations will be apparent to those skilled in
the art. Accordingly, it is intended to embrace all such
alternatives, modifications and variations that fall within the
spirit and broad scope of the appended claims.
All publications, patents and patent applications mentioned in this
specification are herein incorporated in their entirety by
reference into the specification, to the same extent as if each
individual publication, patent or patent application was
specifically and individually indicated to be incorporated herein
by reference. In addition, citation or identification of any
reference in this application shall not be construed as an
admission that such reference is available as prior art to the
present invention. To the extent that section headings are used,
they should not be construed as necessarily limiting.
* * * * *